In fields of medical diagnosis, industrial non-destructive inspection and the like, inspection using an X-ray inspection device such as an X-ray tomograph (hereinafter, described as an X-ray CT scanner) is carried out. The X-ray CT scanner is composed of an X-ray tube (X-ray source) emitting a fan beam X-ray in a fan shape and an X-ray detector including many X-ray detection elements, the X-ray tube and the X-ray detector being arranged opposed each other with a tomographic surface of an inspection target set as a middle. In the X-ray CT scanner, the X-ray tube emits the fan beam X-ray while rotating with respect to the inspection target, and the X-ray detector collects absorption data on X-ray transmitted through the inspection target. Thereafter, the X-ray absorption data is analyzed by a computer, whereby a tomogram is reproduced. For the radiation detector of the X-ray CT scanner, a detection element using a solid scintillator is widely used. In the X-ray detector including the detection element using the solid scintillator, it is easy to increase the number of channels by downsizing the detection element, thus further increasing the resolution of the X-ray CT scanner and the like.
The X-ray inspection device such as the X-ray CT scanner is used in various fields for medical purpose, industrial purpose and so on. As the X-ray CT scanner, for example, there is a known device of a multi-splice type in which detection elements such as photodiodes are vertically and horizontally arranged in two dimensions and a scintillator array is mounted thereon. Employing the multi-splice type makes it possible to superpose cross-sectional images, thereby three-dimensionally expressing the CT image. The X-ray detector mounted in the X-ray inspection device includes detection elements arranged in a plurality of vertical and horizontal lines, and each of the detection elements is provided with a scintillator segment. The X-ray incident on the scintillator segment is converted into visible light, and the detection element converts the visible light into an electric signal to image it. In recent years, to obtain high resolution, the detection element is downsized and the pitch between adjacent detection elements is reduced. Accompanying the above, the size of the scintillator segment is also reduced.
Among the various kinds of scintillator materials used for the above-described scintillator segment, a rare earth oxysulfide-based phosphor ceramics is high in light emission efficiency and has preferable characteristics for use in the scintillator segment. Therefore, an X-ray detector is becoming widely used which is made by combining a ceramic scintillator segment processed by cutout process or grooving process from a sintered compact (ingot) of the rare earth oxysulfide-based phosphor ceramics being the scintillator material and a photodiode as the detection element.
As the scintillator using the phosphor ceramics, there is a known ceramic scintillator composed of a sintered compact of, for example, a gadolinium oxysulfide phosphor. The ceramic scintillator array is fabricated as follows for instance. First, the rare earth oxysulfide-based phosphor powder being the scintillator material is molded into a suitable shape, and the molded powder is sintered into a sintered compact (ingot). The sintered compact of the scintillator material is subjected to a cutting process such as cutout process or grooving process to form scintillator segments corresponding to the plurality of detection elements. A reflective layer is formed between the scintillator segments to integrate them, thereby fabricating the ceramic scintillator array. Further, the scintillator array is required to have a structure that confines light generated by the incident X-ray in the scintillator segments so as to prevent the light from passing through an X-ray incident surface and efficiently takes the light out to the photodiode side. To this end, a reflective layer is formed also on the X-ray incident surface of the ceramic scintillator array.
In the case of using the above-described ceramic scintillator array as the X-ray detector, the dimensional accuracy of the ceramic scintillator array affects the alignment accuracy when bonded to the photodiode and accordingly the resolution of an X-ray CT diagnostic image. Further, a temperature of 50° C. at maximum is applied to the X-ray detector mounted on the X-ray CT scanner. In the scintillator array having the reflective layer containing a resin, expansion of the reflective layer due to heating and contraction due to a decrease in temperature occur to cause a small dimensional change between adjacent scintillator segments, namely, pitch shift of the segment, warpage of the scintillator array, variation in outside dimension and so on. These become a cause of deteriorating the resolution of the diagnostic image of the X-ray detector. In progress of increase in resolution of the diagnostic image of the X-ray detector, a scintillator array having a smaller dimensional change amount due to heating and cooling is required. Further, since the area of the scintillator array also increases with an increase in detection area of the X-ray detector, the control of the dimensional change amount due to temperature change is important. In particular, the warpage of the scintillator array may cause not only a decrease in accuracy due to the dimensional change but also peeling of the reflective layer at the X-ray incident surface.